In radiography, stringent requirements are currently placed on the image quality of the x-ray images. In such images, as are taken especially in medical x-ray diagnosis, an object to be studied is exposed to x-radiation from an approximately point radiation source. Further, the attenuation distribution of the x-radiation is registered two-dimensionally on the opposite side of the object from the x-ray source. Line-by-line acquisition of the x-radiation attenuated by the object can also be carried out, for example in computer tomography systems.
Besides x-ray films and gas detectors, solid-state detectors are being used increasingly as x-ray detectors, these generally having a matricial arrangement of optoelectronic semiconductor components as photoelectric receivers. Each pixel of the x-ray image should ideally correspond to the attenuation of the x-radiation by the object on a straight axis from the point x-ray source to the position on the detector surface corresponding to the pixel. X-rays which strike the x-ray detector from the point x-ray source in a straight line on this axis are referred to as primary beams.
The x-radiation emitted by the x-ray source, however, is scattered in the object owing to inevitable interactions, so that, in addition to the primary beams, the detector also receives scattered beams, so-called secondary beams. These scattered beams, which, depending on the properties of the object, can cause up to 90% or more of the total signal response of an x-ray detector in diagnostic images, constitute an additional noise source and therefore reduce the identifiability of fine contrast differences. This substantial disadvantage of scattered radiation is due to the fact that, owing to the quantum nature of the scattered radiation, a significant additional noise component is induced in the image recording.
In order to reduce the scattered radiation components striking the detectors, so-called antiscatter grids are therefore interposed between the object and the detector. Antiscatter grids include regularly arranged structures that absorb the x-radiation, between which transmission channels or transmission slits for minimally attenuated transmission of the primary radiation are formed. These transmission channels or transmission slits, in the case of focused antiscatter grids, are aligned with the focus of the x-ray tube according to the distance from the point x-ray source, that is to say the distance from the focus. In the case of unfocused antiscatter grids, the transmission channels or transmission slits are oriented perpendicularly to the surface of the antiscatter grid over its entire area. However, this leads to a significant loss of primary radiation at the edges of the image recording, since a sizeable part of the incident primary radiation strikes the absorbing regions of the antiscatter grid at these points.
In order to achieve a high image quality, very stringent requirements are placed on the properties of x-ray antiscatter grids. The scattered beams should, on the one hand, be absorbed as well as possible. On the other hand, the highest possible proportion of primary radiation should be transmitted unattenuated through the antiscatter grid. It is possible to achieve a reduction of the scattered beam component striking the detector surface by a large ratio of the height of the antiscatter grid to the thickness or diameter of the transmission channels or transmission slits, that is to say by a high aspect ratio.
The thickness of the absorbing structure elements or wall elements lying between the transmission channels or transmission slits, however, can lead to image perturbations by absorption of part of the primary radiation. Specifically when solid-state detectors are used, inhomogeneities of the grids, that is to say deviations of the absorbing regions from their ideal position, cause image perturbations by projection of the grids in the x-ray image. For example, in the case of matricially arranged detector elements, there is a risk of projection of the structures of detector elements and antiscatter grids mutually interfering. Perturbing moiré phenomena can thereby arise.
A particular disadvantage of all known antiscatter grids is that the absorbing structure elements cannot be made arbitrarily thinly and precisely, so that a significant part of the primary radiation is always removed by these structure elements.
The same problem occurs in nuclear medicine, especially when using gamma cameras, for example Anger cameras. With this recording technique also, as with x-ray diagnosis, it is necessary to ensure that the fewest possible scattered gamma quanta reach the detector. In contrast to x-ray diagnosis, the radiation source for the gamma quanta lies inside the object in the case of nuclear diagnosis. In this case, the patient is injected with a metabolic preparation labeled with particular unstable nuclides, which then becomes concentrated in a manner specific to the organ. By detecting the decay quanta correspondingly emitted from the body, a picture of the organ is then obtained.
The profile of the activity in the organ as a function of time permits conclusions about its function. In order to obtain an image of the body interior, a collimator that sets the projection direction of the image needs to be placed in front of the gamma detector. In terms of functionality and structure, such a collimator corresponds to the antiscatter grid in x-ray diagnosis. Only the gamma quanta dictated by the preferential direction of the collimator can pass through the collimator, and quanta incident obliquely to it are absorbed in the collimator walls. Because of the higher energy of gamma quanta compared with x-ray quanta, collimators need to be made many times higher than antiscatter grids for x-radiation.
For instance, scattered quanta may be deselected during the image recording by taking only quanta with a particular energy into account in the image. However, each detected scattered quantum entails a dead time in the gamma camera of, for example, one microsecond, during which no further events can be registered. Therefore, if a primary quantum arrives shortly after a scattered quantum has been registered, it cannot be registered and it is lost from the image. Even if a scattered quantum coincides temporally—within certain limits—with a primary quantum, a similar effect arises.
Since the evaluation electronics can then no longer separate the two events, too high an energy will be determined and the event will not be registered. Both situations explain how highly effective scattered beam suppression leads to improved quantum efficiency in nuclear diagnosis as well. As the end result, an improved image quality is thereby achieved for equal dosing of the applied radionuclide or, for equal image quality, a lower radionuclide dose is made possible, so that the patient's beam exposure can be reduced and shorter image recording times can be achieved.
There are currently various techniques for producing antiscatter grids for x-radiation and collimators for gamma radiation. For instance, lamellar antiscatter grids are known, which are made up of lead and paper strips. The lead strips are used for absorption of the secondary radiation, while the paper strips lying between the lead strips form the transmission slits for the primary radiation. However, the limited precision when producing such antiscatter grids, as well as the fact that the thickness of the lead lamellae cannot be reduced further, entail, on the one hand, an undesired loss of primary radiation and, on the other hand, in the case of matricially arranged detector elements of a solid-state detector, problems in the image quality due to moiré stripes and/or grid stripes.
Collimators for gamma cameras are generally produced from mechanically folded lead lamellae. This is a relatively cost-efficient solution, although it has the disadvantage that, in particular when using solid-state cameras with matricially arranged detector elements, for example in the case of cadmium-zinc telluride detectors, perturbing aliasing effects can arise because the structure of these collimators is then relatively coarse.
For producing antiscatter grids for x-radiation, U.S. Pat. No. 5,814,235 A discloses a method in which the antiscatter grid is constructed from individual thin metal film layers. The individual metal film layers include a material that strongly absorbs the x-radiation, and they are photolithographically structured with corresponding transmission holes. To that end, a photoresist needs to be applied on both sides of the respective film and exposed through a photomask.
This is followed by an etching step, in which the transmission holes are etched into the film material. After the remaining photoresist layer has been removed, an adhesion layer is applied to the etched metal films. The metal films are then positioned exactly above one another and are joined together to form the antiscatter grid.
The structure is consolidated by a subsequent heat treatment. In this way, it is possible to produce cellular antiscatter grids with air gaps as transmission channels, which are suitable for applications in mammography and general radiography. In this case, the photolithographic etching technique permits more precise definition of the absorbing and nonabsorbing regions inside the antiscatter grid than is possible with lead lamellae. By using different masks from one metal film to another—in each case with transmission holes that are mutually offset slightly—it is also possible to produce focused antiscatter grids by using this technique. However, an antiscatter grid for x-radiation needs a large number of such metal film layers, which in turn require a large number of different masks and production steps. The method is therefore very time-consuming and cost-intensive.
U.S. Pat. No. 6,185,278 B1 discloses a further method for producing an antiscatter grid for x- and gamma rays, in which individual metal films are likewise photolithographically etched and laminated above one another. In this method, however, in order to produce a focused antiscatter grid, groups of metal film layers with exactly the same arrangement of the transmission holes are assembled together, and only the individual groups have transmission holes arranged mutually offset. This technique reduces the number of photolithographic masks necessary for producing the antiscatter grid.
A further method for producing an antiscatter grid for x-radiation is disclosed by U.S. Pat. No. 5,303,282. This method uses a substrate made of photosensitive material, which is exposed by using a photomask according to the transmission channels to be produced. The channels are then etched from this substrate according to the exposed regions. The surface of the substrate, as well as the inner walls of the transmission channels, are coated with a sufficient thickness of a material that absorbs the x-radiation. In order to increase the aspect ratio, a plurality of such prepared substrates are optionally stacked above one another. Similar production techniques for producing cellular antiscatter grids for x-radiation are described in EP 0 681 736 B1 or U.S. Pat. No. 5,970,118 A. Etching transmission channels into thicker substrates, however, leads to a loss of precision of the channel geometry.
The publication by G. A. Kastis et al., “A Small-Animal Gamma-Ray Imager Using a CdZnTe Pixel Array and a High Resolution Parallel Hole Collimator” discloses a method for producing a cellularly constructed collimator for gamma radiation. In this case as well, the collimator is produced from laminated layers of metal films, here made of tungsten, which are photochemically etched. This production method is therefore also very elaborate and cost-intensive.
Post-published DE 101 47 947 describes a method for producing an antiscatter grid or collimator using the technique of rapid prototyping. In this method, the geometry of the transmissive and the nontransmissive regions of the antiscatter grid or collimator is set first. Next, by way of a rapid prototyping technique through layer-wise solidification of a structural material under the action of radiation, a base body is constructed according to the geometry of the transmissive regions, and is coated with a material which strongly absorbs x or gamma radiation on the inner surfaces of the transmission channels formed and on the front and rear surfaces. The layer thickness is selected in this case such that incident secondary radiation is virtually completely absorbed in this layer.
By using a rapid prototyping technique when constructing the base body, very filigree structures can be produced with very high accuracy. The base body can be produced very simply in this way, without needing to perform a large number of elaborate method steps. In this method, the structures, particularly the intermediate walls or webs, forming the absorbing regions, between the transmission channels can be realized in a simple way with a thickness of approximately 60–200 μm. The production of intermediate walls with thicknesses below 60 μm continues, however, to require substantial outlay. On the other hand, absorbing intermediate walls with a thickness of 60–200 μm lead to an unfavorable primary beam transparency of the antiscatter grid or collimator.